Ex vivo remodeling of excised blood vessels for vascular grafts

ABSTRACT

The present invention provides an ex vivo vascular remodeling methods and system by which an excised, small diameter blood vessel can be harvested and expanded to provide viable vascular grafts, as demonstrated at the physical and molecular levels, and as optimized in vivo. The tissue-engineered vessels generated by the present invention closely resemble native vessels in terms of structure, histologically, including endothelial coverage and intricate structural components such as the internal elastic lamina, viability (as measured with the MTT assay and TUNEL analysis), and function (vasoactivity, mechanical and biomechanical properties). Thus, the resulting vascular grafts behave in a manner similar to native arteries in terms of mechanical integrity, and provide clinically relevant patency rates when implanted in vivo. Moreover, the ex vivo methods and system permit the precise control of the mechanical environment involving the excised vessel, while at the same time permitting carefully monitoring of the resulting growth/remodeling, thereby opening new avenues of research regarding the mechanical stimuli responsible for specific aspects of remodeling in vivo.

REFERENCE TO RELATED APPLICATIONS

This application is a continuation of application Ser. No. 10/165,461,filed Jun. 7, 2002, herein incorporated by reference in its entirety,which claims priority to 60/297,203, filed Jun. 8, 2001, hereinincorporated by reference in its entirety.

GOVERNMENT INTERESTS

This invention was supported in part by the National Institutes ofHealth Grant No. R01 HL64388-01A1. The Government may have certainrights in this invention.

FIELD OF THE INVENTION

This invention relates generally to the field of tissue remodeling,specifically the ex vivo remodeling of blood vessels for use as vasculargrafts.

BACKGROUND OF THE INVENTION

More than a century ago, based on observations of the microvasculature,Thoma proposed that longitudinal tension controls vessel length. Sincethen, a number of studies have shown that blood vessels can remodeleither physiologically or pathologically when exposed to alteredmechanical environments. Arteries exposed to elevated flow (such asarteries upstream of arteriovenous fistulas (Holman, Surgery 26:889-917(1949); Shenk et al., Surg. Gynecol. Obstet. 110:44-50 (1960)),collateral arteries carrying flow around an obstruction (Mulvihill, etal., N. Engl. J. Med. 104:1032 (1931)), and aortorenal bypass grafts(Stanley, et al., Surgery 74:931 (1973)) remodel (autoregulate) toincrease their luminal diameter in response to increased flow as theresult of vascular smooth muscle cell relaxation. In contrast, arteriesexperiencing reduced flow decrease luminal diameter.

Animal studies substantiate these clinical observations and suggest thatvessels remodel so as to restore the wall shear stress to initial levels(Fung et al., J. Appl. Physiol. 70(6):2455-2470 (1991); Kamiya et al.,Am. J. Physiol. 239(1):H14-21 (1980); Zarins J. Vasc. Surg. 5(3):413-420(1987)). Inflation of a tissue expander implanted within a rat hind limbover different periods of time ranging from 2 to 21 days increased thelength of adjacent blood vessels 83±43%. Relatively slowexpansion-induced lengthening (≦10% per day) did not diminish vesselpatency, though more rapid expansion did substantially reduced patency(Stark, Plastic and Reconstructive Surgery, 30(4):570-578 (1986)).

However, the complex interdependence between components of themechanical environment (e.g., pressure, shear, and strain) in vivo hashindered the identification of the specific mechanical stimuliresponsible for remodeling. For example, by altering the viscosity ofthe perfusing medium, Melkumyants and coworkers have reported that bydecoupling the effects associated with shear rate, ∂v_(z)/∂r, (e.g.,convection-enhanced transport and streaming potentials) and the wallshear stress, −μ∂v_(z)/∂r, that acute autoregulation is a response towall shear stress, not to flow rate per se (Melkumyants et al.,Cardiovasc Res. 24(2): 165-168 (1990)). Several widely used systems thatexpose cultured endothelial and smooth muscle cells to well-definedmechanical environments exist, but extrapolating results from cellculture models to vascular remodeling has proven to be problematic.

Traditional organ culture models employing excised vessels, such ashuman saphenous veins under static conditions, provide a well-definedchemical/biochemical environment and have been used to study the effectsof pre-existing intimal hyperplasia, surgical preparations (Soyomo etal., Cardiovasc. Res. 27(11):1961-1967 (1993)), and specific biochemicalfactors, including bFGF (Soyomo et al., 1993) and ET-1 (Porter et al.,J. Vasc. Surg. 28(4):695-701 (1998); Masood, et al. Brit. J. Surg.84(4):499-503 (1997) on intimal hyperplasia. The inadequacy of thesemodels is evidenced by the fact that vessels maintained under staticconditions, even in the absence of known biochemical atherogenicstimuli, rapidly undergo pathological remodeling, including substantialintimal hyperplasia (Soyomo et al., 1993).

The atherogenic nature of traditional organ culture models appears to beat least partially due to the absence of physiologically relevant levelsof mechanical forces. Porter and coworkers developed a crude,first-generation flow system by cutting an excised saphenous veinlongitudinally and gluing the adventitial surface of the vein to theinside of a perfused Tygon tube (Porter et al., Cardiovasc. Res.31(4):607-614 (1996)). The application of venous levels of pressure andflow-induced shear stress to excised human saphenous veins partiallyattenuated intimal hyperplasia associated with traditional organculture, while arterial levels of pressure and shear stress completelyabolished intimal hyperplasia (Porter et al., 1996). These resultsshowed that, with a mechanically active environment, it was possible tomaintain blood vessels in organ culture for weeks without pathologicalchanges.

While the mechanical environments used in these studies were intended tomimic aspects of the arterial or venous circulation, they lacked manyrelevant mechanical features, including temporal variations, cyclicstrains, as well as pressure drops across the vessel wall and theresulting transmural flow—each of which is a potentially importantmechanical stimulus to blood vessels as summarized in reviews by, e.g.,Gooch et al., Mechanical Forces: Their Effects on Cells and Tissues,Berlin, Springer, 182 (1997), and by Liu, Crit. Rev. Biomed. Eng.27(1-2):75-148 (1999). Perfusion systems have been developed and used toprovide a sophisticated mechanical environment by introducing pulsatileflow, cyclic flexure (Vorp et al., Ann. Biomed. Eng. 27(3):366-371(1999)) and transmural pressure (Chesler et al., Am. J. Physiol. 277(5Pt 2):H2002-2009 (1999)). These have been used to study the effects ofthe mechanical environment on gene expression (Vorp et al, 1999),endothelial cytoskeleton (Herman et al., J. Cell Biol. 105(1):291-302(1987), lipid transport across the endothelium (Herman et al, 1987), andvasomotor responses (Labadie et al., Am. J. Physiol. 270(2 Pt2):H760-768 (1996)).

Perfusion systems have also been used to investigate the effect ofhydrodynamic forces on endothelial cells, with specific focus on themechanisms by which endothelial cells perceive a mechanical stimulus andconvert it to the initial biochemical response (i.e.,mechanotransduction) (Gooch et al., Am. J. Physiol. 270(2 Pt 1):C546-51(1996)), as well as the effect of biochemical pathways stimulated byfluid flow and mechanical forces on cellular proliferation (Gooch etal., J. Cell Physiol. 171(3):252-258 (1997); Gooch et al., MechanicalForces: Their Effects on Cells and Tissues, 1997)) and susceptibility toviral infection. In addition, the effect of a hydrodynamic environmenton the development of tissue-engineered cartilage has been investigated(Gooch, K., et al., “Mechanical Forces and Growth Factors,” in Frontiersin Tissue Engineering, (C. Patrick, A. Mikos, and L. McIntire, editors.)Pergamon, N.Y. p. 61-82 (1998)).

Vessel cultures have also been used to explore the molecular biology ofvascular remodeling, both under static (Porter et al., 1998; Masood etal., 1997; Porter et al., Brit. J. Surg. 85(10):1373-1377 (1998); Porteret al., Eur. J. Vasc. Endovasc. Surg. 17(5):404-412 (1999)), andmechanically active environments (Chesler et al., 1999; Meng et al.,1999). One area in which the ex vivo vessel models have beenparticularly insightful is mechanical regulation of matrixmetalloproteinases (MMPs), expression and activity (Vorp et al, 1999;Chesler et al., 1999; Meng et al., Exp. Mol. Pathol. 66(3):227-237(1999); Mavromatis et al., Arterioscler. Thromb. Vasc. Biol. 20(8):1889-1895 (2000)), and the role of MMPs in vascular remodeling (Porteret al., 1998; Porter et al., 1999; Loftus et al., Ann. N Y Acad. Sci.878:547-50 (1999)).

Tenascin-C (TN-C) is large (>1000 kDa), disulfide-linked, hexamericextracellular matrix (ECM) glycoprotein that is prominently expressedduring embryonic development, epithelial-mesenchymal interactions, woundhealing, cancer, and notably, vascular disease (Mackie, Int. J. Biochem.Cell Biol. 29(10): 1133-1137 (1997)), and is also subject to mechanicalregulation. TN-C expression has been shown to be increased in rats andchildren suffering from pulmonary hypertension (Jones et al., J. CellSci. 1 12(Pt 4):435-445 (1999)), and under increased mechanical loadingregimes, TN-C expression co-localizes with neointimal lesions expressingepidermal growth factor (EGF) and proliferating cell nuclear antigen(PCNA) (Jones et al., J. Cell Biol. 139(1):279-293 (1997); Jones et al.,Circ. Res. 79(6):1131-1142 (1996)). The pro-proliferative role of TN-Cis supported by in vitro studies that show TN-C acts as a survivalfactor for cultured smooth muscle cells (Cowan et al., Circ. Res.84(10):1223-1233 (1999)). The majority of studies show that soluble,extracellular, and matrix factors regulate TN-C at the transcriptionallevel (Chiquet-Ehrismann et al., Bioessays 17(10):873-878 (1995)). Inaddition, targeted suppression of TN-C arrests progressive pulmonaryhypertrophy in organ culture (Cowan et al., 1999). Taken together, thesedata strongly suggest that in the vessel wall the expression of TN-C isregulated by the mechanical environment, and the expression of thisprotein in turn is a key regulator of SMC proliferation and vascularremodeling.

Nevertheless, there is a sizable unmet demand for effectivesmall-diameter vascular prostheses for use in coronary bypass surgery.Currently, the best replacements for occluded arteries are autologousarteries, which have a cumulative patency rate of 93% after 5 years(Lytle et al., J. Thorac. Cardiovasc. Surg. 89(2):248-258 (1985)).However, the number of expendable autologous arteries of appropriatedimensions for bypass grafts is severely limited, although there arenumerous expendable arteries of smaller dimensions.

In animal studies where autologous tissue-engineered small-diametervessels were evaluated in vivo, they performed much worse than anautologous vein would have (e.g., about half of the tissue-engineeredvessels had decreased perfusion or loss of patency within 1 month(Niklason et al., Science 284(5413):489-493 (1999); Campbell et al.,Circ. Res. 85(12): 1173-1178 (1999)). Donor veins of appropriatedimensions are more readily available and are frequently used, but theyhave a substantially lower patency. Human saphenous vein grafts have apatency of ˜90% at early time points, and 81 % after 1 year (Fitzgibbonet al., J. Am. Coll. Cardiol. 28(3):616-626 (1996)), but this has beenreported to diminish to 45% after 5 years (Lytle et al., 1985).

Thus, the limited availability of suitable autologous arteries, coupledwith the poor long-term patency of autologous veins, has led researchersto explore a number of approaches to create small-diameter vascularprostheses. These include using natural (Sandusky et al., J. Surg. Res.58(4):415-420 (1995)) and synthetic polymeric materials (Smith et al.,J. Med. Chem. 39(5):1148-1156 (1996); Uretzky et al., J. Thorac.Cardiovasc. Surg. 100(5):769-776 (1990)), pre-endothelializing existingtypes of polymer grafts in vitro (Stansby et al., Cardiovasc. Surg.2(5):543-548 (1994); Stansby et al., Brit. J. Surg. 81(9):1286-1289(1994)), and creating bioartificial or tissue-engineered blood vesselsfrom cells and various support structures (Weinberg et al., Science231:397-400 (1986); L'Heureux et al., J. Vasc. Surg. 17(3):499-509(1993); L'Heureux et al., FASEB J. 12(1):47-56 (1998); Tranquillo et al.Biomaterials 17(3):349-357 (1996); Niklason et al., 1999); Shinoka etal., J. Thorac. Cardiovasc. Surg. 1 15(3):536-545 (1998)). While thereare a number of different approaches to generating autologoustissue-engineered vessels in vitro, they all follow the same generalparadigm: isolate specific cell types from blood vessels, expand thesecells in vitro, and reassemble these cells into a tissue-engineeredblood vessel—with the last step being the major challenge.

Many of these approaches yielded tissue-engineered arteries that grosslyresemble native vessels, but in animal studies where tissue-engineeredvessels generated in vitro were evaluated in vivo, their performance wasinferior to that of autologous veins (Niklason et al., 1999; Campbell etal., 1999; Fitzgibbon et al., 1996). However, it was generally foundthat the performance of autologous blood vessels (whole vessels) wasclearly superior to that of tissue-engineered blood vessels (preparedfrom only cells derived from the vessels).

There remains, however, a need in the art for a method or system bywhich a blood vessel can be harvested and used to direct the growth ofan intact vessel ex vivo, wherein the newly formed vessel would be ofsufficient size to permit the formation of a tissue-engineered vessel,which would be suitable for use as an arterial graft in vivo. Criteriafor assessing the remodeled arteries relate both to the extent that thevessels grow ex vivo, and the degree that the remodeled arteriesresemble healthy arteries of corresponding dimensions. Even modestincreases in vessel dimensions would be potentially useful. Based onrough estimates using Poiseuille's law (i.e., vessels deformiso-volumetrically (Milnor, Hemodynamics, 2^(nd), 1989)), increasing theinternal arterial diameter by 33% will increase the ability of thatartery to carry blood by more than 200%. Poiseulle's law,${Q = {{- \frac{\pi\quad\Delta\quad P}{8\quad\mu\quad L}} \cdot r^{4}}},$relates the volumetric flow rate, Q, to the radius of a straightcylindrical tube of radius, r. Increasing the radius from 100% X to 133%X, increases flow from Y to 3.1 Y, a greater than 200% increase in flow.

In addition, in light of the foregoing and because blood vessels in vivoactively remodel (i.e., change size and/or composition) in response tochronic changes in the mechanical environment, the utilization of thisability of intact blood vessels to remodel supports the use of thesystem and methods of the present invention as a more effective andalternate approach to generating tissue-engineered blood vessels.

SUMMARY OF THE INVENTION

The present invention provides a system and method by which a smallblood vessel is harvested with minimal morbidity of the donor, and thediameter, length, and wall thickness of the excised vessel are increasedby subjecting the vessel to the appropriate mechanical environment exvivo over time. Thus, a tissue-engineered vessel is produced, which issuitable for use as a blood vessel graft in vivo.

Clinical observations and animal studies indicate that vessels remodelin response to altered mechanical environments, but the complexinterdependence between components of the mechanical environment (e.g.,pressure, shear, and strain) in vivo has hindered the identification ofthe specific mechanical stimuli responsible for specific aspects ofremodeling. To identify the mechanical stimuli responsible for vascularremodeling, an ex vivo perfusion system is provided for exposing viable,excised blood vessels, cells and tissues to precisely controlled flowand pressure regimes, while maintaining the viability of the vessel. Theexcised vessels are housed in a medium-filled chamber, cannulated oneach end, and perfused with cell culture medium supplemented with serumand antibiotics (FIG. 1). As in traditional cell or organ culturesystems, temperature, pH, pO₂, pCO₂, and nutrient composition areregulated.

In addition, the system allows for the control of several key aspects ofthe mechanical environment. It is an advantage of this system over theprior art to offer improved control of the mechanical environment andallow real time monitoring of vessel remodeling. An understanding ofwhich mechanical stimuli control vascular remodeling is utilized torationally direct the remodeling of vessels ex vivo.

In a preferred embodiment of the invention the ex vivo perfusion systemwas used to determine which aspects of the mechanical environment directthe remodeling of arteries. Moreover, experimental data havedemonstrated the ability to control and measure extravascular pressurein accordance with the provided methods.

Furthermore, based on preliminary data, it appears reasonable to expectincreases of vessel length of at least 100%. Unless the length andinternal diameter of a vessel are greatly increased, little to no medialthickening maybe required. For example, rough estimates based onLaplace's Law suggest that if the internal diameter of the vesselincrease 33%, and thickness of the vessel remains constant, the stressesin the arterial wall will increase a proportional 33%, which is arelatively small increase. This assumes that the wall thickness is smallcompared to the vessel diameter, and that the average stress across thewall thickness can be estimated with Laplace's equation, which statesthat the transmural pressure, p_(i)-p_(o)=T(l/r). Therefore, the hoopstress, T/h, is directly proportional to the radius.

Using carotid arteries as a model system, data is presenteddemonstrating that mechanically induced, directed remodeling of excisedarteries is possible and that the structure and function of theresulting arteries remain comparable to native arteries. At theconclusion of each experiment, vessels were harvested and processed forhistology. From histological sections and subsequent immunostaining,indices of vascular remodeling (intimal, medial, and adventitialthickness) and injury (proliferation of, extracellular matrix synthesisby, and phenotypic change of vascular smooth cells, disruption ofinternal elastic lamina, formation of a neointima, loss of endothelium)were quantified. By subjecting excised porcine arteries to well-definedmechanical environments, it is shown that, in arteries, transmuralpressure drop regulates wall thickness, longitudinal tension regulateslength, and flow-induced shear stress regulates inner diameter.

Thus, it is an object of this invention to provide a system and methodby which small blood vessels, such as arteries, or even veins, can beharvested with minimal donor site morbidity and remodeled ex vivo,thereby engineering blood vessels for autologous small-caliber vasculargrafts. In addition, the ex vivo system is advantageously applied tobetter understand the molecular biology of vascular remodeling byfacilitating the testing of hypotheses not amenable to study using invivo or cell culture models.

It is a further object of this invention to explore the extent to whicharteries can be elongated ex vivo, to determine the identity of themechanical factors that regulate arterial lumenal diameter and wallthickness, and to explore the extent to which the lumenal diameter andwall thickness of arteries can be increased ex vivo. In vivo studieswill evaluate the efficacy of arteries elongated ex vivo as autologousarterial graphs in model subjects, to provide data for eventual humanapplication.

It is also an object of the invention to utilize the ex vivo perfusionsystem to explore the molecular regulation of mechanically inducedvascular remodeling by characterizing the expression and regulation ofkey regulator factors. For example, the spatial expression anddistribution of TN-C mRNA and protein resulting from various mechanicalloads is monitored in the cultured vessels to determine the region(s) ofthe TN-C promoter responsible for mechanosensitivity.

Additional objects, advantages and novel features of the invention willbe set forth in part in the description, examples and figures whichfollow, and in part will become apparent to those skilled in the art onexamination of the following, or may be learned by practice of theinvention.

DESCRIPTION OF THE DRAWINGS

The foregoing summary, as well as the following detailed description ofthe invention, will be better understood when read in conjunction withthe appended drawings. For the purpose of illustrating the invention,there are shown in the drawings, certain embodiment(s) which arepresently preferred. It should be understood, however, that theinvention is not limited to the precise arrangements andinstrumentalities shown.

FIGS. 1A and 1B diagrammatically depict the existing ex vivo perfusionsystem. FIG. 1A is a schematic diagram of an embodiment using only onevessel. FIG. 1B shows an enlargement of the chamber housing the bloodvessel 5 from FIG. 1A.

FIG. 2 graphically illustrates volumetric flow rate (upper line) andpressure (lower red line) vs. time for an excised porcine carotid arteryexposed to a mechanical environment intended to simulate its nativearterial environment. Before harvesting the vessel, the averagevolumetric flow rate was 320 ml/min as measured using transit-timeultrasound.

FIG. 3 graphically compares artery length for neonatal elongationexperiments (n=5). Each artery length was normalized by individual exvivo unloaded length (i.e., all arteries unloaded lengths are 1 on day0); average data are shown with SEM. (**) indicates p<0.005.

FIGS. 4A-4D photographically depict the immunohistochemistry of paraffinsections prepared from porcine carotid arteries before (FIGS. 4A and 4B)and after (FIGS. 4C and 4D) 10 days of perfusion. Staining for smoothmuscle α actin (FIGS. 4A and 4C) and for elastin (FIGS. 4B and 4D)allows identification of the internal elastic lamina (IEL) and themedia, both of which are important landmarks used to quantify vascularremodeling.

FIGS. 5A and 5B graphically depict as a function of time, the effect ofcontrolling extravascular pressure to atmospheric pressure (FIG. 5A,bottom line, lightest gray), to a fixed amount above atmosphericpressure (FIG. 5A, top two lines, black and dark gray), or such that thecalculated transmural pressure is a constant (FIG. 5B, bottom curve). InFIG. 5B, intravascular pressure is shown in the top curve (black line),and extravascular pressure is shown in the middle curve (gray line).

FIGS. 6A-6F photographically depict a microscopic assessment of arteriescultured ex vivo for 9 days under a mechanically active environment. InFIG. 6A, hematoxylin and eosin staining reveal healthy vessel. In FIG.6B, immunostaining for the smooth muscle cell specific isoform ofα-actin strongly stains the media, but not the adventia. In FIG. 6C,immunostaining is specific for proliferating cell nuclear antigen(PCNA). FIG. 6D is a scanning electron micrograph of the luminal face ofthe vessel. In FIG. 6E, immunostaining of the extracellular matrixprotein elastin reveals the internal elastic lamina and underlyingstriations. In FIG. 6F, terminal dUTP nick-end labeling reveals a verylow rate of apoptosis/necrosis.

FIGS. 7A and 7B graphically depict artery length for juvenile elongation(n=6, FIG. 7A) and control (n=4, FIG. 7B) experiments. Each arterylength was normalized by individual ex vivo unloaded length (i.e., allartery unloaded lengths are 1 on day 0); average data are shown withSEM. Perfused refers to loaded length of the arteries while in theperfusion system, while the unloaded length refers to arteries out ofthe system, under no load. (**) indicates p<0.005.

FIGS. 8A-8F provide a comparison of representative histological sectionsof fresh and elongated arteries from juvenile pigs taken at 10×.Sections were stained with hematoxylin and eosin (H & E, FIGS. 8A and8B), PCNA (FIGS. 8C and 8D), and the TUNEL assay (FIGS. 8E and 8F).Arrows in FIGS. 8C, 8D, 8E and 8F indicate approximate lumen location(lumen always faces right).

FIGS. 9A and 9B graphically depict an assessment of arterial function ofporcine carotid arteries cultured for 9 days ex vivo. In FIG. 9A, theeffect of adding a KCl solution to the medium bathing the artery isrecorded as the time average of the pressure over 1 second. In FIG. 9B,the data shown represents the mechanical testing of a strip cut from anartery on an Instron machine.

FIG. 10 graphically depicts the average longitudinal stress-strainrelationship for fresh (n=9), elongated (n=4), and control (n=5)arteries from juvenile pigs. Data are shown until the first point offailure for each group (e.g., the minimum ultimate strain from the freshgroup was 65%). Data were unavailable below 65% for control arteries.

FIG. 11 is a comparative line graph depicting the effect of ex vivoremodeling on the longitudinal length of a vessel after 9 days. Thelength of the vessel shown by line A (the in vivo length) was equal tothat in line C (the initial length in the system). Line B shows thefreshly excised length showing elastic recoil from the in vivo length.Line D was the length of the vessel after 9 days in the ex vivo system(showing a 100% increase over the in vivo length. Line E shows thelength of the vessel after removal from the ex vivo system, which evenafter recoil was 70±3% greater than the initial length. The length ofline D=2.0× that of line C. The length of line E=1.7× that of line B.Conditions B and E are unstressed (i.e., no applied longitudinal loads);while all other conditions involve longitudinal loading.

FIG. 12 graphically depicts outer vessel diameter as a function of timeof a porcine artery exposed to pulsatile flow.

FIG. 13 graphically depicts pressure harmonics for the pressure-vs.-timecurve shown in FIG. 2. The pressure harmonics are derived from thecoefficients for the Fourier series. Consistent with the observations ofothers, the coefficients rapidly approach zero with increasing harmonicnumber; and truncating the Fourier series at n=10 accurately replicatesthe observed pressure versus time.

DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION

The invention provides a system and method by which appropriatemechanical environments are applied ex vivo to direct the remodeling ofsmall, excised blood vessels to create tissue-engineered vesselscharacterized by increased length, internal diameter, and wallthickness. Thus, the small excised vessels, arteries, or even veins,become tissue-engineered blood vessels for use in vascular surgery. Theinvention further provides an evaluation of the performance of thesetissue-engineered blood vessels in vivo.

The disclosed ex vivo system allows investigations of the hypothesisthat longitudinal stress or strain induces artery elongation. Inaddition, while there are autologous donor arteries with proper diameterand wall thickness for vascular grafts, they often are of aninsufficient length to meet the required need. For example, the internalthoracic artery has excellent long-term patency, but is of an adequatelength for only a single bypass graft. However, recognizing that if theartery could be elongated, it could be used to bypass multipleocclusions, and the use of vessels demonstrating inferior performancecould be avoided, the present invention advantageously provides reliabletissue-engineered blood vessels of sufficient length to meet this need.In addition, the ex vivo perfusion system is further used to explore themolecular regulation of mechanically induced vascular remodeling bycharacterizing the expression and regulation of key regulatory factors,for which the spatial expression and distribution of mRNA and proteinare monitored as a result of various mechanical loads.

Thus, the invention provides a protocol by which localized intravascularand extravascular pressures are measured in real time, and the measuredpressures are compared with the calculated pressure estimates. In apreferred embodiment of the invention the ex vivo perfusion system wasused to determine which aspects of the mechanical environment direct theremodeling of arteries. Moreover, experimental data have demonstratedthe ability to control and measure extravascular pressure in accordancewith the provided methods.

The Ex Vivo Perfusion System

The ex vivo perfusion system (FIG. 1A) was designed and built with acapability of independent control specific aspects of the mechanicalenvironment (e.g., the magnitude and time rate of change inintravascular pressure and flow). Turning to FIG. 1A, a HarvardApparatus pulsatile blood pump 1 pushes fluid clockwise around thecircuit. If desired, the initial pulsatile pressure/flow profiles aredampened by the compliance chamber 2, wherein the extent of dampeningdependent on the volume of gas present. Pressure in the chamber housingthe excised blood vessel 5 is controlled by needle valves 3, which areup and downstream of the chamber. Pressure 4 and low 6 are measured 250times per second with in-line probes attached to the correspondingTriton Cardiovascular Measurement Modules (models 200-204 and 200-206)(8 and 9, respectively). Analogue output from these modules is digitizedand sent to a personal computer for analysis and storage using LabView.

Medium is pooled in a reservoir 7, which permits gas exchange, before itis returned to the pump. The system is enclosed in a 37° C. environment.

For ease of presentation, only one vessel is shown in the embodimentpresented in FIG. 1A. However, multiple vessels can be run in parallel,each having its own housing and, when necessary, correspondingcompliance chambers and needle valves.

FIG. 1B provides an enlarged diagram of the chamber housing the bloodvessel 5 from FIG. 1A. Vessels are cannulated with two slidingstainless-steel tubes and the entire assembly is inserted into thePlexiglas cylinder. The stainless steel tubes slide independently of therest of the unit to control the vessel strain. Ports on the Plexiglascylinder allow the entry and exit of bathing medium and blood gasmixture.

The prototype perfusion system, however, was limited in its ability tocontrol other aspects of the mechanical environment that may beimportant for vascular remodeling. Therefore, studies were undertaken toexpand the capabilities of the system for exposing excised blood vesselsto well-defined mechanical environments to enable a) improved control ofthe mechanical environment and b) real-time monitoring of vascularremodeling. The resulting instrument system has been improved to nowenable real-time measurement of pressure and volumetric flow.

Elementary instrumentation principles were employed to interfacecommercial flow and pressure meters with the PC used for data analysisand storage. Basic fluid mechanics concepts (e.g., those embodied in theNavier-Stoke's equation and Poiseuille equation) were used to guide thedesign of the existing system capable of independent control ofpressure. To ensure that the flow through the vessel is fully developed,the length of straight constant-diameter tubing immediately upstream ofthe vessel is greater than that calculated for fully developed steadyand unsteady flow, 27 cm and 11 cm respectively.

The inlet length for steady flow, L_(s), (the length required toestablish a velocity profile with deviation of less than 1% fromparabolic) is given by the equation L_(s)=0.16 rNR_(R), where N_(R) isthe Reynolds number. This relationship holds when N_(R)≧50 (Fung,Biomechanics: Circulation, 2nd ed. (1996), herein incorporated byreference). N_(R) for this system is 675, when the Reynolds number isdefined as N_(Re)=ρrU/μ, where ρ is the density of the fluid and μ isthe dynamic viscosity. The viscosity of medium is approximately that ofwater or about ⅕^(th) of that of blood. The average velocity, U, iscalculated from U=Q/πr_(i) ².

The unsteady entry length, L_(us), is approximated by the equation,L_(us)=2.64 U/ω, where ω is the pulse rate in radians (Fung, 1996). Anunderstanding of the effects of pulsatility on the velocity profile wasaided by considerations of the Womersley number. For example, for thehydrodynamic conditions of the porcine carotid artery, the WomersleyNumber, α, is 1.1.α=2r(ω/v)^(1/2)where ω is the pulse rate in radians, r is internal radius of thevessel, and v is the kinematic viscosity (Fung, 1996). The relativelylow Womersley number indicates that the transient inertial force is ofthe same order as the shear force, suggesting that the pulsatile flowcan be crudely approximated as a parabolic velocity profile.

The system was validated by running excised vessels (saphenous vein,jugular vein, and carotid artery) in the perfusion system for up to ˜10days. FIG. 2 illustrates representative flow and pressure recordingsfrom an excised porcine carotid artery exposed to a mechanicalenvironment intended to simulate its native arterial environment.Vessels exposed to ex vivo culture were subsequently characterized withimmunohistochemistry performed on paraffin sections (FIGS. 4A-4D).

Improved System Offers Enhanced Control of Mechanical Environment.

Extravascular pressure: In the embodied perfusion system, theextravascular pressure (i.e., the pressure inside the chamber housingthe excised blood vessel) is maintained at 0 mm Hg gauge (i.e., ambientatmospheric pressure) by allowing the chamber to vent to the atmospherethrough a 0.22 μm filter. This situation is an appropriate model of invivo conditions where extravascular pressure is roughly atmosphericpressure, but it limits investigations into the role of pressures invascular remodeling. As a result, the perfusion system was modified toallow for control of extravascular pressure: 1) at a given constantlevel, or 2) to provide constant transmural pressure.

Constant extravascular pressures are provided by attaching to thechamber housing the excised blood vessel a side arm with a fixed heightof medium exposed to the atmosphere at the top surface. Theextravascular pressure, P_(o), is estimated from hydrostatics asP_(o)=ρgz, where ρ is the density of the fluid (1.03 g/cm³), g is theacceleration due to gravity, and z is the height of the column of fluid.The validity of this approach has been demonstrated in preliminarystudies, the results of which are shown in FIGS. 4A-4D.

An alternative to maintaining a given extravascular pressure, involvesfixing transmural pressure. To do so, the intravascular andextravascular pressures are first set while under no flow conditions(e.g., by using fixed heights of medium exposed to the atmosphere at thetop surface). The forces exerted by these two pressures are balanced bythe tension generated in the vessel wall. To a first approximation, thecircumferential tension generated by the vessel, T, can be calculatedfrom the law of Laplace for thin walls, T=ΔPr, as the wall thickness his much less than the vessel diameter, r.

Radial strain: The outer diameter of the vessels is measured in realtime using a laser scanning system (Model # LX2-V10W from Keyence,Woodcliff Lake, N.Y.) (FIG. 12). The embodied system is capable ofnon-invasively measuring over 250 times per second in vessel diametersup to 1 cm, with a repeatability of 5 μm. Since the measurement deviceis external to the perfusion system, with only the scanning laser beamentering, it is easy to relocate the device to measure diameters atdifferent points and in different regions along the vessel (e.g., nearthe end and at the endpoint). Radial strain, ε_(θ), is calculated asε_(θ)=(D−Do)/Do, where D is the outer diameter of the blood vessel at agiven time and Do is the initial value.

Longitudinal strain: The embodied perfusion system allows for changingthe longitudinal strain of the vessel by sliding (extending) thestainless steel tubes on which the vessel is mounted (FIG. 1B). Sincethe length of the vessel is only being changed very slowly (e.g., inseveral mm steps once per day), the length is determined by manuallymeasuring the position of the stainless steel tubes entering the vesselchamber. Sonomicrometry is also used to measure the distance between twopoints on the same side of the outer surface of the vessel. These pointswill be approximately 2 cm apart, and located along the middle portionof the vessel. From these points, longitudinal strain, ε_(z), iscalculated as ε_(z)=(L−Lo)/Lo, where L is the length of the vessel at agiven time, and Lo is the initial length. The perfusion system allowsfor changing the longitudinal strain of the vessel by sliding thestainless steel tube on which the vessel is mounted.

The capabilities of the proposed system, wherein the mechanicalenvironment is controlled by the proposed perfusion system device, aresummarized in Table 1. TABLE 1 Key aspects of mechanical environmentcontrolled by perfusion system. Mechanical parameter Magnitude Time rateof change Flow rate (Q) 0-500 ml/min Steady to ˜3 Hz cycles Pulse rate0-200 beats/min Intravascular pressure (P_(i)) 0-500 mm Hg Steady to ˜3Hz cycles Extravascular pressure (P_(o)) 0-500 mm Hg Steady or in-phasewith intravascular pressure Transmural pressure (P_(i)-P_(o)) 0-500 mmHg Steady or in-phase with intravascular pressure Longitudinal strain(ε_(z)) 0 to 100% Steady to very slow changes (e.g. 10%/day max.) Radialstrain (ε_(θ)) 0 to 10% Steady to ˜2 Hz cycles

All pressures are gauge pressures.

The extravascular pressure is controlled to atmospheric pressure (FIG.5A, light gray curve, lowest line), to a fixed amount above atmosphericpressure (FIG. 5A, black and gray curves, remaining lower lines), orsuch that the transmural pressure is a constant (FIG. 5B, bottom curve).

Next the extravascular compartment is sealed to create a constant volumeextravascular system, thus fixing the vessel radius. More precisely,∫₀^(l)π  r²  𝕕r,where l is the length of the vessel, must be a constant volume at anytime. The assumption that the vessel radius is fixed at all times is onecase, but definitely not the only case that satisfies this integral. Thevelocity at which the pressure wave travels along the length of thevessel (typically several meters per second) is assumed to be rapidcompared to the radial motion of the vessel. However, fixing the radius(and therefore the transmural pressure) for all times is a solution.

The validity of the analysis is confirmed by measuring transmuralpressure under the test conditions. Because the pressure drop across thevessel wall depends on the vessel radius and the material properties ofthe vessel wall, to the extent that neither of these parameters changes,the transmural pressure is regulated. Acknowledging that theseassumptions are not trivial, the validity of the conclusions wereevaluated by measuring the transmural pressure across ex vivo vessels inreal time.

Real-Time Monitoring of Vascular Remodeling

Geometric Remodeling: In the original perfusion system, vascularremodeling could only be assessed at the conclusion of the study whenthe vessel was fixed and histological sections were prepared, however bythe present invention quantification of vascular remodeling has beenimproved by allowing real-time monitoring of vessel diameter. Althoughlaser-scanning techniques are used to measure radial strain, innerdiameter (i.e., lumenal diameter) cannot be measured by this method.Therefore, echo-ultrasound may be used to estimate wall thickness, h.The internal diameter, D_(i), is calculated as D_(i)=D_(o)−2 h.

Biomechanical Remodeling: Two independent procedures are used todetermine the viscoelastic biomechanical properties of the arteriesduring ex vivo culture: pressure-wave propagation analysis andpressure-diameter analysis (Milnor, Hemodynamics, 2nd ed., Baltimore,Williams & Wilkins (1989)). The primary benefit of these two techniquesis that they permit continuous real-time evaluation of the mechanicalproperties of the vessel wall throughout the ex vivo culture period.

“Pressure-wave propagation analysis” compares the pressure harmonicsresulting from Fourier transformation of the pressure-vs-time profilegenerated at several points along the vessel to determine the waveattenuation coefficient, a (which is related to the viscous nature ofthe vessel) and the true phase velocity, c (which can be used tocalculate the dynamic elastic modulus of the vessel). The three-pressuremethod of wave propagation analysis, which mathematically removes theeffects of wave reflection (Gessener et al., IEEE Trans. Biomed. Eng.13:2-10 (1996), herein incorporated by reference), is used, asimplemented by others to assess vascular remodeling in vivo (Wells etal., Am. J. Physiol. (Heart Cir. Physiol.) 274:H1749-H1760 (1998)).

Consistent with the observations of others, the coefficients rapidlyapproach zero with increasing harmonic number and truncating the Fourierseries at n=10 accurately replicates the observed pressure versus timedata and the use of additional terms (e.g., n up to 30) did notnoticeably improve accuracy (data not shown).

P(t) is measured at 3 equally distant points within the vessel usingcatheter pressure transducers. A Fourier analysis is performed on eachpressure profile and the resulting harmonics expressed in complex form(FIG. 13). The coefficients for the Fourier series are as follows:${P(t)} = {a_{0} + {\sum\limits_{n = 1}^{\infty}\left( {{a_{n}{\cos\left( {n\quad\pi\quad{t/T}} \right)}} + {b_{n}{\sin\left( {n\quad\pi\quad{t/T}} \right)}}} \right)}}$were solved by numerically integrating the integrals$a_{n} = {\frac{1}{T/2}{\int_{{- T}/2}^{T/2}{{P(t)}{\cos\left( {n\quad\pi\quad{t/T}} \right)}\quad{\mathbb{d}t}}}}$for n 0, 1, 2, . . . and$b_{n} = {\frac{1}{T/2}{\int_{{- T}/2}^{T/2}{{P(t)}{\sin\left( {n\quad\pi\quad{t/T}} \right)}\quad{\mathbb{d}t}}}}$for n=1, 2, 3 . . . . The pressure harmonic, ΔP, and the phase angle, φ,were then obtained using the relationships M=√{square root over (a²+b²)}and φ=arctan(b/a). Alternatively P(t) can be expressed as a complexnumber in the form${P(t)} = {\frac{A_{0}}{2} + {\frac{1}{2}{\sum\limits_{n = 1}^{\infty}{\left( {A_{n} - {j\quad B_{n}}} \right){\mathbb{e}}^{j\quad n\quad\omega\quad t}}}} + {\frac{1}{2}{\sum\limits_{n = 1}^{\infty}{\left( {A_{n} = {j\quad B_{n}}} \right){{\mathbb{e}}^{{- j}\quad n\quad\omega\quad t}.}}}}}$

Thus, the resulting harmonics are substituted into Bergel's equation forthe true wave propagation coefficient,$\gamma = {\frac{1}{\Delta\quad x}{\cosh^{- 1}\left( \frac{P_{1} + P_{3}}{2P_{2}} \right)}}$

The true wave propagation coefficient describes the transmissioncharacteristics of each pressure harmonic as it travels through anartery. It consists of a real portion, which is the attenuationcoefficient a, and an imaginary portion, which is the angular frequencydivided by the true phase velocity, c (i.e., γ=a+(ω/c). The dynamicelastic modulus, E_(dyn), is related to true phase velocity by theequation E_(dyn)=3 ρr_(o)/(h(2-h/r_(o)))c² where ρ is the density of thecell culture medium, h is the arterial wall thickness, and r_(o) is theexternal diameter.

In addition, “pressure-diameter transient analysis” harmonics resultingfrom Fourier transformation of the pressure and external radiustransients over the pulse cycle is used to calculate the complexviscoelastic modulus (E*) using the equation of Bergel (J. Physiol.(Lond) 156:458-469 (1961)), which is herein incorporated by reference,$E^{*} = {\left\lbrack {\frac{3r_{i}^{2}r_{o}}{2\left( {r_{o}^{2} - r_{i}^{2}} \right)} \cdot \frac{M}{\Delta\quad r_{o}}} \right\rbrack \cdot {\mathbb{e}}^{({j\quad\theta})}}$where r_(i) and r_(o) are the internal and external radii, respectively,M is the amplitude of the pressure harmonic, Δr_(o) is the amplitude ofthe radius harmonic, θ is the phase angle between the correspondingpressure and radius harmonics, and j is √{square root over (−1)}. As atest for internal consistency, the real component E* from thepressure-diameter transient analysis is compared to the dynamic elasticmodulus, E_(dyn), obtained form the pressure wave propagation analysis.

To further assess the accuracy of the real-time measurements ofmechanical properties performed while vessels are in the ex vivoperfusion system, static and dynamic stress-strain relationships aremeasured from axial and longitudinal strips prepared from selectvessels. The static and dynamic stress-strain measurements are made on,e.g., a fully digital Instron machine (model 5543) with a positionalaccuracy of 0.156 μm (FIG. 9B).

In addition to facilitating the determination of the applied forces thatmodulate remodeling (e.g., absolute pressure or transmural pressure),the ex vivo perfusion system provides insight into the actual stressesto which the vessels actually respond. By measuring the acute variationsin vessel diameter in response to cyclic changes in measured transmuralpressure, it is possible to estimate some of the stresses in the vesselwall. Following the Berceli analysis of the biomechanics of excisedarteries (Brant et al., J. Biomechanics 21(2):107-113 (1988)) and veins(Berceli et al., J. Biomech. 23(10):985-989 (1990)) (each of which areincorporated by reference) exposed to various hemodynamic conditions,the axial stress (T_(zz)) and hoop stress (T_(θθ)) are estimated.

Each of these parameters can be expressed as functions of theincremental modulus (E_(inc)), essentially the real component of thecomplex viscoelastic modulus applied over a limited range of strain.This is calculated as follows:${E_{inc} = {\frac{{TP}_{\max} - {TP}_{\min}}{r_{o,\max} - r_{o,\min}} \cdot \frac{2\left( {1 - \sigma^{2}} \right)r_{i,{avg}}^{2}r_{o,{avg}}}{r_{o,{avg}}^{2} - r_{i,{avg}}^{2}}}};$${T_{zz} = {\frac{\sigma\quad E_{inc}}{\left( {1 - \sigma^{2}} \right)} \cdot \frac{\eta}{r_{i,\min}}}};$$T_{\theta\quad\theta} = {{\frac{E_{inc}}{\left( {1 - \sigma^{2}} \right)} \cdot \frac{\eta}{r_{i,\min}}} + \frac{E_{inc}h^{2}\eta}{12{r_{i,\min}^{3}\left( {1 - \sigma^{2}} \right)}}}$where TP is transmural pressure, r is radius, μ is dynamic fluidviscosity, h is wall thickness, and η is the measured displacement ofthe vessel wall, and the subscripts min, max, and avg refer to theminimal (diastolic) value, the maximal (systolic) value, and averagevalues, respectively.

To a very close approximation, Poisson's ratio, σ, is 0.5 for bloodvessels (i.e., vessels deform iso-volumetrically). In this case, thevessel wall is considered elastic, axisymmetric, semi-infinite inlength, straight with circular cross-section, constrained from motionlongitudinally and the radial displacement is small compared to theradius. These calculations provide estimates of the mechanical stressesin the vessel wall and the incremental modulus aid in the quantificationof remodeling. The calculated mechanical stresses are correlated withthe observed vascular remodeling.

EXAMPLES

The invention is further described by example. The examples, however,are provided for purposes of illustration to those skilled in the art,and are not intended to be limiting. Moreover, the examples are not tobe construed as limiting the scope of the appended claims. Thus, theinvention should in no way be construed as being limited to thefollowing examples, but rather, should be construed to encompass any andall variations which become evident as a result of the teaching providedherein.

Although the disclosed experiments were conducted using porcine vesselsas models to allow for the detailed in vivo evaluation of thetissue-engineered vessels, the findings are directly applicable to humanvascular replacement and provide the foundation for human tissuestudies.

For all experiments, vessels were harvested from anesthetized pigs priorto euthanization. Using aseptic technique, an incision were made, thevessel were separated from surrounding fascia and connective tissue, andthe vessel was excised. The vessel was briefly washed in buffer andsubmerged in cell culture medium until placed in the perfusion system nomore than 2 hours later. Unless stated otherwise, the duration of eachexperiment was 4 weeks, which has been shown to be an adequate amount oftime to observe substantial vascular remodeling in vivo.

Example 1 The Perfusion System: Control of Mechanical Environment

The perfusion system consisted of a peristaltic pump, compliancechamber, artery chamber, and reservoir, all connected using Tygonlaboratory tubing (Formula R-3603, Fisher Scientific, Pittsburgh, Pa.),ports for injection into or sampling from the perfusing medium, andpressure transducers (Model MER100, Triton Technology, Inc., San Diego,Calif.) upstream and downstream from the artery (FIGS. 1A and 1B).Steady flow was provided by a Masterflex roller pump (1) (Model 7553-70,Cole-Parmer, Vernon Hills, Ill.) with Masterflex Tygon LFL pump tubing(Formula 06429-25, Fisher Scientific, Pittsburgh, Pa.). Real-timepressure data were acquired via an analog-digital board (ModelPCI-6023E, National Instruments, Austin, Tex.) connected to a TritonSystem 6 Twinpak Chassis (Active Redirection Transit-Time Flow Module,Model 200-206 and Dual Pressure Amplifier Module, Model 200-204, Triton,San Diego, Calif.). Data were visualized and recorded using aLabView-based routine (LabView Full Development System, NationalInstruments, Austin, Tex.) on a PC. Gas exchange was provided to boththe artery chamber and reservoir via 5% CO₂ bubbling chambers. Theentire system, except for the roller pump, was maintained in a dark, 37°C. environment. All components were sterilized using ethylene oxide andassembled under sterile conditions.

The pulsatile-flow pump forces medium through the ex vivo perfusionsystem with a well-defined volumetric flow rate. Controlling thecompliance and the resistance of the system allows for a wide range ofmechanical environments (with respect to magnitude and time rate ofchange) ranging from arterial to venous conditions as well as supra- andsub-physiological conditions. As shown in FIG. 2, the measured ex vivopressure (lower line) and volumetric flow profiles (upper line) aremaintained in a mechanical environment at values that simulate typicalconditions of a porcine carotid artery in vivo. Typical hemodynamicvalues for pigs are a pulse of ˜80 beats per minute and arterial bloodpressure ˜100/60 mm Hg. Before harvesting the vessel, the averagevolumetric flow rate was 320 ml/min as measured using transit-timeultrasound. In addition to replicating in vivo pressure and flow profilequalitatively, specific quantitative features were also accuratelyreproduced (FIG. 2).

Table 2 summarizes key aspects of the mechanical environment controlledby the existing perfusion system and the ranges over which theseparameters can be controlled. TABLE 2 Mechanical parameter MagnitudeTime rate of change Flow rate 0-500 ml/min Steady to ˜3 Hz Intravascularpressure 0-500 mm Hg Steady to ˜3 Hz Pulse rate 0-200 beats/min Pulsepressure 0-500 mm Hg Steady to ˜3 Hz Extravascular pressure   0 mm HgSteady Longitudinal strain 0 to 100% Steady to ˜10%/day

Note that while some of these parameters are independent of one another(e.g., average intravascular pressure and average flow can beindependently controlled), other parameters are coupled (temporalvariations in pressure and flow are linked). Though it would be ideal tohave independent control of each mechanical parameter, this is notalways feasible. For example, radial strain is dependent on parametersthat can be directly controlled (e.g., intravascular and extravascularpressure), as well as other parameters that cannot be directlycontrolled (e.g., wall thickness and mechanical properties of thevessel, such as modulus). Therefore, given the number of degrees offreedom in the system, it is not possible to arbitrarily set Pi, Po andε_(o).

Example 2 Determining Which Mechanical Factors Regulate Remodeling ofArteries

Artery Harvest, Preparation and Maintenance: Carotid arteries fromneonatal (˜5-kg) and juvenile (˜30-kg) pigs were harvested bycardiothoracic surgeons at the Children's Hospital of Philadelphia afterthe animals were euthanized. Carotid arteries from adult pigs (˜100-kg)were obtained from freshly exsanguinated pigs at a local abattoir.Arteries were transported in ice-cold culture medium (Dulbecco'sModified Eagle's Medium (DMEM) supplemented with 10% fetal bovine serum,100 U/mL penicillin and 100 μg/mL streptomycin, all from LifeTechnologies, Inc., Rockville, Md.). Upon arrival, arteries wereprepared within a laminar flow hood using sterile instruments.

Arteries, measuring 3-6 cm in length, were individually cleaned ofexcess adventitial and connective tissue. Sections were taken forhistology, methylthiazol tetrazolium (MTT) assay and, in some cases, dryweight and/or mechanical testing. Dry weight measurements were madeafter at least 8 hours in a Speedvac system (SC100, Thermo-Savant,Holbrook, N.Y.).

Arteries were individually installed into the artery chamber andpressurized with medium to locate leaks. Installation consisted ofcannulating the artery onto ribbed stainless steel rods (stainless steel8, 10 or 13 gauge microtubing, McMaster-Carr, Dayton, N.J.) via silksutures, where the outer diameter of the rod roughly matched the innerdiameter of the artery. Whole, leak-free artery segments were installedat the approximated in vivo length prior to perfusion (initial ex vivoloaded length) unless stated otherwise. The initial extension ratio (exvivo loaded to unloaded length) was determined for each artery fromneonatal and juvenile animals by measuring the length of the arterybefore and after harvest (unloaded). For arteries from adult animalswhere in vivo length was not measurable, a ratio of 1.5 was used, sincethe average ratio from neonatal and juvenile arteries was 1.47±0.03.

After installation of the artery, the chamber was filled with ˜200 mL of37° C. culture medium, completely submerging the artery. The chamber wasthen connected to the perfusion system containing ˜500 mL of 37° C.culture medium, wherein the desired volumetric flow rate had beenpreviously established (10-15 mL/min). Steady flow was then diverted tothe artery chamber from the bypass branch.

Carotid Arteries from Neonatal Pigs: Five carotid arteries obtained fromneonatal pigs were installed in the ex vivo perfusion system at theirphysiological loaded length and elongated ⅙^(th) of the initial loadedlength (16.7%/day) on days 2 to 7 of a 9 experiment. Control arteriesfrom neonatal pigs were cultured under identical conditions at fixedlength (n=6, separate study by (Clerin et al., Ann. Biomed. Eng.29(suppl):S-145 (2001)). All arteries were perfused at the approximatedin vivo volumetric flow rate of 50 ml/min.

For the higher volumetric flow rates (50 mL/min, arteries from neonatalpigs) the flow rate was increased slowly over a 2-hour period until thedesired flow rate was achieved. The flow rate for neonatal arteries waschosen to approximate the in vivo flow rate for neonatal carotidarteries, however subsequent studies determined that subphysiologicalflow rates were necessary to abrogate de-endothelialization and massivecell death in neonatal arteries (Clerin et al., 2001).

Upon removal from the perfusion system, control arteries retained noincrease in unloaded length, while elongated arteries retained a65.2±4.5% increase in unloaded length (FIG. 3). Under these flowconditions, histological evaluation revealed that both elongated andcontrol arteries were denuded of their endothelial cells, had lost mostof their cellularity, especially in the inner medial region, and hadhigh levels of cell death. The average MTT index was 0.35±0.13 (n=5)demonstrating low viability as compared to control arteries whichmeasured 0.87±0.26 (n=6, p=0.06).

Carotid Arteries from Juvenile Pigs: Because of these findings, the flowrate for juvenile and adult arteries were both chosen to be 10-15mL/min, 5-10% of the approximated in vivo flow rates for each artery.The artery chamber was maintained at atmospheric pressure by venting thechamber to the atmosphere via a 0.22-μm filter. The average time fromharvest to installation in the perfusion system was 60-90 minutes forneonatal and juvenile and 2-3 hours for adult arteries.

A total of 18 carotid arteries from juvenile pigs were perfused in theex vivo perfusion system and either elongated (n=12, “elongatedarteries”) or held at physiological loaded length (n=6, “controlarteries”) for 9 days (Table 3). All juvenile arteries were installed attheir physiological loaded length and perfused at a volumetric flow rateof 10-15 ml/min, previously shown to be within the optimum range forretaining artery viability (Clerin et al., 2001). TABLE 3 Summary of exvivo culture experiments. Age of Volumetric Donor Flow Rate Length Pigs(ml/min) Longitudinal Strain Protocol n increase Rupture Neonatal 50100% increase in physiological loaded 5  5** 0 length in 9 days (16.7%on days 2 to 7) Juvenile 10-15 Fixed physiological loaded length, 9 days6 0 0 15 50% increase in physiological loaded 8  6** 2 length in 9 days(8.3% on days 2 to 7) 15 66% increase in physiological loaded 1 0 1length in 7 days (13.2% on days 2 to 6) 15 100% increase inphysiological loaded 3 0 3 length in 9 days (16.7% on days 2 to 7) Adult10 100% increase in physiological unloaded 1 0 0 length in 7 days (16.7%on days 2 to 7) 10 100% increase in physiological unloaded 2 0 0 lengthin 27 days (5% on days 4 to 23) 10 100% increase in physiological loaded1 0 1 length in 9 days (16.7% on days 2 to 7) 10 100% increase inphysiological loaded 2 0 2 length in 27 days (5% on days 4 to 23)A significant increase in unloaded length (p < 0.005) is denoted by(**).

Carotid Arteries from Adult Pigs: All carotid arteries from adult pigswere subjected to rapid protocols (stretched 16.7% on days 2 to 7 of a 9day experiment, n=1) or slow stretching protocols (stretched 5% on days4 to 23 of a 27 day experiment, n=3), ruptured prior to completion ondays 6, 15 (n=2), and 19 (Table 3). Rupture was avoided by installingarteries at ex vivo unloaded length, and elongating 5% of the unloadedlength on days 4 to 23 of a 27-day experiment (n=2), or elongating 16.7%of the unloaded length on days 2 to 7 of a 9-day experiment (n=1,removed on day 7 due to suture failure).

None of the arteries retained an unloaded length increase upon removalfrom the perfusion system. Arteries installed at unloaded length allincreased their wet weight (41.6±1.7%, n=3) while those that rupturedshowed no clear trend (8.5±13.7%, n=3). Viability of all arteries, asassessed by MTT index, was similar to fresh arteries.

Application of Longitudinal Strain (Elongation Protocol): In preliminarystudies, it was demonstrated that by applying a longitudinal strain,vessels could be elongated ˜100% over 9 days. Longitudinal strain wasapplied daily to the artery by manual displacement of the steel rods.Arteries perfused for 9 days were held at their physiological length(initial loaded length) on day 1, stretched at a rate of ⅙^(th) or1/12^(th) of the physiological length per day from day 2 to 7, held atthe final stretched length on day 8, and excised on day 9. Similarly,arteries perfused for 27 days were held at constant length (initialloaded or unloaded length) on days 1 to 3, stretched 1/20^(th) of theinstalled length on days 4 to 23, held at the final stretched length ondays 24 to 26, and excised on day 27. Control arteries were cultured inthe perfusion system under identical conditions, but were held at theirphysiological length (initial loaded length).

The in vivo length of the porcine arteries were noted prior to excision.The length of the arteries without an applied load were measured and thevessels were placed in the perfusion system at their in vivo lengths.Vessels were randomly assigned to two groups. Vessels in the first group(control) were arteries were cultured in the perfusion system underidentical conditions, but were held at their physiological length(initial loaded length). Vessels in the second group were subjected todifferent longitudinal strain rates (˜2 to 20% per day) for the variousindicated durations (1 week to 2 months). At the end of the experiment,the lengths of the vessels without applied loads were measured andcompared. All unloaded lengths reported were measured at least 15minutes after removal from the artery chamber since no significantchange in artery length (>0.1 mm) was seen after this time.

Wall thickness: Increased mean intravascular pressure (i.e.,hypertension) results in remodeling of blood vessels characterized byincreased ratio thickness and alterations in the zero-stress state ofthe vessel (Fung et al., 1991). In vivo, hypertension results inincreased transmural pressure that in turn results in increased radialstrain and transmural flow making it difficult to identify whichmechanical stimulus is responsible for the observed remodeling. Ex vivo,it is possible to independently control intravascular and extravascularpressure.

To determine whether intravascular pressure was affecting remodelingdirectly, or whether it was acting by its effect on transmural pressure,porcine carotid arteries were exposed to the three sets of conditionssummarized in Table 4. Vessels in the control group were exposed tonormal arterial pressures; vessels in the “normal” hypertension groupwere subjected to elevated intravascular pressure, but a normalextravascular pressure, as is normally the case with hypertension. Boththe intravascular and extravascular pressures were increased by an equalamount so that transmural pressure remains normal for vessels in the“corrected” hypertension group. The fact that vessels in the normalhypertension group, but not the other two groups, experience medialthickening was a confirmation that intravascular pressure affectsremodeling by its effect on transmural pressure, as opposed to directly.TABLE 4 A summary of experimental groups. Number IntravascularExtravascular Transmural Condition of vessels pressure pressurepressure 1) Control 3 100/60  0 Normal 2) Normal 3 200/160 0 Elevated   hypertension 3) “Corrected” 3 200/160 100 Normal    hypertension

All pressures are in mm of Hg gauge. Flow is pulsatile with ˜1 Hz cycleand mean volumetric flow of 300 ml/min.

Internal diameter: To evaluate the hypothesis that chronic changes inluminal diameter resulting from vascular remodeling are also dependenton the wall shear stress, chronic studies were conducted, similar to theacute studies of Melkumyants et al., 1990. The viscosity of theperfusion medium was varied from 1 to 10 cP by the addition of highmolecular weight dextran, a compound that is not harmful to excisedvessels in chronic cultures (Chesler et al., 1990). Excised porcinecarotid arteries were perfused under the conditions described in Table4. A summary is presented in Table 5 of the experimental groups used toinvestigate the relative contribution of fluid flow and shear stress onvascular remodeling leading to increases in internal diameter. TABLE 5Number of Flow rate Viscosity Initial shear Condition vessels (ml/min)(cP) stress (dyn/cm²) 1) Control 3 300 5 Normal 2) High flow/ 3 1500 1Normal    normal shear 3) High flow/ 3 1500 5 5x normal    high shear 4)Normal flow/ 3 300 25 5x normal    high shear

Po=100 mm Hg for all conditions.

In all conditions, the flow is steady. Therefore, flow can be consideredas fully developed laminar flow in a circular conduit of constant crosssection where the wall shear stress, τ_(rz), is calculated as:τ_(rz)=4 Qμ/πr_(i) ³

The internal diameter of the vessels was assessed throughout theexperiment. At the conclusion of the experiment, the arteries were fixedat a pressure of 100 mm Hg and histological sections are prepared. Thefact that shear stress regulates chronic changes in lumenal diameter isshown by the finding that groups 1 and 2 have the same diameter as eachother, but they have a smaller diameter than that which was found ingroups 3 and 4.

Results: The maximum elongation while retaining mechanical integrity andviability was achieved by stretching 1/12^(th) of the physiologicalloaded length (8.3%) on days 2 to 7 of a 9-day experiment. Six arterieswere successfully lengthened in the perfusion system 48.1±2.8% from theinitial physiological loaded length (p<0.001). The correspondingincrease from initial to final ex vivo unloaded length upon removal fromthe system was 20.5±3.3% (p<0.005)(FIG. 7A).

In contrast, none of the six control arteries perfused at physiologicalloaded length for 9 days (n=6) retained a length increase upon removalfrom the perfusion system (FIG. 7B).

The wall thickness of control arteries was significantly lower than bothelongated (p<0.005) and freshly harvested arteries (p<0.005), whereasthe wall thickness of elongated arteries was similar to freshlyharvested specimens (Table 6). TABLE 6 Material properties for juvenilearteries. Freshly Harvested Elongated Control Wall thickness (mm) 0.80 ±0.04 (n = 9) 0.87 ± 0.08 (n = 6) 0.48 ± 0.02 (n = 5)**,++ Change in wetweight (%) N/A 39.9 ± 18.4 (n = 5) 21.5 ± 2.0 (n = 2) Dry/wet weight (%)13.1 ± 0.9 (n = 6)  12.2 ± 0.4 (n = 4)  14.7 ± 2.8 (n = 2)Data are shown ± standard error of the mean (SEM).Significant differences were found between fresh and control arteries(**), and elongated and control arteries (++), with p < 0.005.

As compared to freshly harvested arteries, the wet weight of elongatedarteries increased 39.9±18.4% (n=5, p=0.07), whereas the wet weight ofcontrol arteries increased 21.5±2.0% (n=2, p=0.06) (Table 6). Thedry/wet weight ratio was not significantly different between fresh,control or elongated arteries (Table 6).

Three arteries elongated 1/12^(th) of their physiological length on days2 to 7 were removed before day 9. One was removed from the perfusionsystem on day 7 (after an elongation of 50%) due to a slow leak, butshowed no other problems and was included in the analysis.

More rapid elongation (i.e., >10%/day) always caused arteries fromjuvenile animals to rupture (Table 3). Three arteries ruptured whenelongated ⅙^(th) of the physiological loaded length daily (16.7%/day) ondays 2, 3 and 5 of 9, while one artery failed on day 3 of 7 whenelongated ⅛^(th) of the physiological loaded length daily (12.5%/day) ondays 2 to 6. None of the arteries that ruptured retained a permanentunloaded length increase upon removal from the system.

While both elongation of the juvenile vessels in the perfusion system(while under load) and the increased unloaded length indicate arterialelongation, the more relevant parameter is the increase in length atphysiological longitudinal stress. Noting that the average physiologicallongitudinal strain is 50% (bar 1 vs. bar 2 of FIG. 7A), the averagephysiological longitudinal stress at 50% strain for fresh arteries is0.40 MPa (mega Pascals)(FIG. 8).

The longitudinal strain of elongated arteries at 0.40 MPa is 72%. Takentogether (i.e., the product of the unstressed length and thelongitudinal strain at physiological stress), these data indicate thatarteries from juvenile pigs elongated for 9 days ex vivo are 40% longerthan equivalent fresh arteries at physiological longitudinal stresses.Interestingly, the stress-strain curve for elongated arteries was foundto closely resemble the curve for control arteries. Thus, the increasein longitudinal extensibility at relatively low stress appears to be aresult of ex vivo culture, rather than the increase in the appliedlongitudinal stress or strain.

Several non-exclusive mechanisms may contribute to the observedelongation of juvenile arteries including plastic deformation due to theapplied longitudinal stress/strain (i.e., creep), mechanically-induced,biologically-mediated redistribution of tissue components (i.e.,remodeling without growth), and mechanically-induced,biologically-mediated deposition of new tissue components (i.e.,growth). While both creep and remodeling without growth could accountfor limited lengthening of the arteries, substantial elongation ofarteries without substantially decreasing wall thickness or innerdiameter would require growth as well. As a result, the 40% increase inwet weight of the arteries as the result of the 9-day elongation processshows that growth is occurring during the elongation of the arteries.Associated with this increase in wet weight, there is a small(1.07-fold) increase in hydration of the arteries, but the majority ofincrease in wet weight is due to the 29% increase in the dry weight ofthe elongated arteries.

The greater ability of juvenile arteries to remodel as compared to adultarteries is consistent with data from in vivo studies showing that bothadult and juvenile arteries can remodel in response to changes in theirmechanical environment, but that juvenile respond more readily. Langilleet al., Am. J. Physiol. 256:H931-939 (1989); Miyashiro et al., Circ.Res. 81:311-319 (1997)). By comparison, neonatal arteries elongated upto 100% under load and 65% when unloaded within 9 days, though asreported by Clerin et al., 2001, even control neonatal arteries hadreduced viability over 9 days in culture with physiological flow rates.

Example 3 Control of Extravascular Pressure

Since vessels are compliant viscoelastic materials, adequate control ofthe extravascular pressure was essential to validate the accuracy of theestimates for transmural pressure. Accordingly, to critically contolextravascular pressure (i.e., the pressure of the medium bathing theexternal surface of the vessel), pressures were measured by placing acatheter pressure transducer close to the external surface of a segmentof compliant Penrose tubing used as a surrogate of an artery for thesepreliminary studies and compared measured values to the set pointvalues.

A goal of this experiment was to reduce both the magnitude and thevariation in the transmural pressure across a compliant tube (e.g.,30±10 mm Hg controlled, as opposed to 120±25 mm Hg uncontrolled). Asshown in FIG. 3B, these data indicate that extravascular pressure wasaccurately controlled to a constant value, and transmural pressure wascontrolled with a degree of success which appeared to be sufficient toevaluate the relative role of transmural pressure in vascularremodeling.

These findings are adaptable to studies with arteries because whenmeasured pressures are found to vary significantly from predictedvalues, the measured values were used for subsequent analyses.Therefore, the degree of control of transmural pressure, which wasobtained when supplemented with measurements of the intravascular andextravascular pressures, enabled detailed study of the effects of therelative contribution of absolute pressure, transmural pressure, andcyclic strain on vascular remodeling, each of which is an aspect of themechanical environment affecting vascular cells and blood vessels.

Example 4 Viability, Structure, and Function of Arteries after Ex VivoCulture

Porcine carotid arteries were harvested and cultured in the ex vivoperfusion system under mechanical active environments for up to 9 daysas described in Example 2. Vessels were harvested at select times (timezero, 1 hr and 1, 3, 5, and 9 days) and the viability, structure, andfunction of the vessels were assessed using various criteria summarizedin Table 7. The following results compare freshly harvested porcinecarotid arteries to vessels perfused ex vivo. TABLE 7 Assay Measure ofMajor results/implications Viability MTT Mitochondrial activityViability not diminished after 9 days in culture PCNA # of proliferatingcells in histological Increased cell proliferation section through-outfull thickness of vessel wall TUNEL # of cells with fragmented DNA inNormal levels of apoptosis and histological section (indicating necrosisapoptosis or necrosis) Structure H&E Histological/microscopic structureGeneral arterial structure preserved; No intimal hyperplasia ElastinStains internal elastic lamina (IEL) in Internal elastic lamina andhistological sections elastic layer in media intact Smooth Stains smoothmuscle cells in Strong staining in media similar muscle histologicalsections to fresh isolated arteries α actin SEM Microstructure ofluminal surface Endothelium intact, but with some cells rounded upFunction Macroscopic Occlusion, aneurysms, and Vessels intact andnon-occluded Assessment “hemorrhage” Addition of Vasoactive response ofvessel Voltage-gated calcium channels KCl exposed to contractilestimulus and contractile apparatus indicated by altered pressure dropfunctional along vesselAbbreviations:MTT—3(4,5-dimethylthia-zolyl-2)-2,5-dihenyl tetrazolium bromide;PCNA—proliferating cell nuclear antigen;TUNEL—terminal pUTP nick-end labeling;H&E—hematoxylin and eosin;SEM—scanning electron microscope.

By all criteria employed so far, ex vivo cultured vessels have beenshown to be nearly identical to freshly harvested vessels.

Histology and TUNEL Assay: Ring samples (˜1 mm in length) were takenfrom fresh and cultured arteries for TUNEL assays and histologicalevaluation. Samples were fixed overnight in either 70% ethanol or 10%formalin, dehydrated, embedded in paraffin, and cut into 5 μm thicksections, which were mounted onto glass slides. Slides weredeparaffinized and stained with hematoxylin and eosin (H & E), the PC10antibody recognizing proliferating cell nuclear antigen (PCNA/HRP, DAKO,Carpinteria, Calif.), and the in situ Cell Death Detection kit, POD(TUNEL, Roche Molecular Biochemicals, Indianapolis, Ind.) according tomanufacturer's instructions or common protocols.

Mitochondrial activity was assessed using the methylthiazol tetrazolium(MTT) assay (Sigma, St. Louis, Mo.). Artery ring samples approximately1-2 mm in length were incubated in 0.5 mL of 1 mg/mL MTT solution for 24hours at 37° C., rinsed with 0.9% saline solution, cut into 2-5 pieces,placed in covered containers containing 5 ml of isopropanol andincubated at room temperature for at least 24 hours. The absorbance of 1mL of the liquid was measured at 550 nm and normalized by the dry weightof the sample. An MTT index was defined as the final normalized MTTvalue divided by the initial normalized MTT value. An index value near 1indicated that mitochondrial activity was similar for fresh and culturedspecimens.

Viability index measured by MTT was 0.34±0.03 units/mg before perfusionand 0.30±0.08 units/mg after 9 days of perfusion (n=6, p=0.66); thefrequency of cells containing fragmented DNA, as measured by the TUNELassay, were low in both sets of vessel indicating little apoptosis ornecrosis (FIG. 5F); and the rate of cell division, indicated by thepresence of proliferating cell nuclear antigen (PCNA), was slightlyhigher in the culture arteries than the freshly harvested arteries (FIG.5C). The elevated proliferation was not the result of intimalhyperplasia. The gross macroscopic and microscopic structure of thearteries was not changed by ex vivo culture (FIGS. 6A and 6D), nor wasthe tissue-specific localization of ECM (FIG. 6E) or cells altered (FIG.6B). Cultured arteries continued to exhibit vasoactive responsiveness(FIG. 6A).

The cellularity, structure, and viability of freshly harvested, control,and successfully (i.e., not ruptured) elongated arteries were similar,as assessed in histological sections stained with H & E (FIGS. 8A and8B), PCNA (FIGS. 8C and 8D) and TUNEL (FIGS. 8E and 8F). There was noevidence of intimal hyperplasia or fragmentation of the internal elasticlamina in any of the elongated or fresh arteries (FIGS. 8A and 8B). Fourof the elongated arteries and five of the control arteries had goodendothelial coverage, while two elongated and one control artery weredenuded of their endothelial cells. One of the denuded elongatedarteries had been denuded (for unknown reasons) prior to installationinto the perfusion system, and thus had no endothelial cells afterelongation.

TUNEL staining of fresh, elongated and control samples revealed minimalcell death (FIGS. 8E and 8F). An exception to this finding was that thetwo elongated arteries and one control artery which were denuded ofendothelial cells stained strongly for TUNEL, consistent with previousfindings that denuded arteries experience progressive cell deathbeginning in the inner lumen by day 9, irrespective of elongationprocedures (Clerin et al., 2001).

Test for Vasoactivity: Vasoactivity experiments were performed on selectarteries from juvenile animals upon completion of elongation protocols.Pressure was measured upstream and downstream of the artery in real timeto yield the pressure drop caused by the change in arterial innerdiameter as endothelial independent vasoactive agents were added to theartery chamber.

Porcine carotid arteries were cultured for 9 days ex vivo. A KClsolution was added to the medium bathing the artery (FIG. 1B) causingthe vessel to contract as indicated by the increased average pressureupstream (i.e., an increased pressure drop along the length of thevessel). Though instantaneous pressure fluctuated, as shown in FIG. 2,the time average of the pressure over 1 second is displayed in FIG. 9A.The data represents the response in one vessel, but similar responseswere observed in the other vessels investigated.

Vasoconstriction of cultured arteries in response to KCl (FIG. 9A) is asalient observation, because tissue-engineered vessels generated by themethods of L'Heureux et al. and Niklason et al. do not respond to thisstimuli (Nicholas L'Heureux, personal communication), indicating thatthe smooth muscle cells have lost important aspects of their basicfunction (e.g., their voltage-gated calcium channels), perhaps as theresult of their expansion in two-dimensional cell culture prior to theiruse to form engineered vessels.

Immediately prior to removal from the perfusion system on day 9, threeelongated arteries were tested for vasoactive response to norepinephrine(NE) and sodium nitroprusside (SNP), a NO donor. NE was added to theartery chamber to cause vasoconstriction. The arteries were thenmonitored for roughly 60 minutes, at which time SNP was added to theartery chamber to cause dilation. The pressure drop across the arterieswas measured over time to monitor the constriction and dilation causedby these agents.

All arteries tested contracted in response to NE (1×10⁻⁶ M (n=2) or1×10⁻⁴ M (n=1)), which caused a decrease in lumen diameter correspondingto a peak pressure increase of 52.7±30.3 mm Hg in an average time of14.8±4.1 minutes. Addition of SNP (1×10⁻⁴ M, n=3) caused an increase inlumen diameter corresponding to an average pressure decrease of38.6±16.0 mm Hg in an average time of 6.8±0.4 minutes.

Evaluation of Mechanical Properties: Samples from select freshlyharvested and cultured arteries were evaluated for mechanical propertiesas follows. Arterial sections approximately 1-2 cm in length weretransported in room temperature medium and cut into sheets by onelongitudinal incision. Throughout all mechanical evaluation, thespecimens were kept at room temperature and constantly hydrated withcalcium-free phosphate buffered saline (PBS). The thickness of thespecimen was measured at 3 locations using a near frictionless LVDTprobe and platform apparatus, after the probe was allowed to reach anequilibrium value (60 seconds).

A “dogbone” stamp was used to cut out a representative sample from thesheet aligned in the longitudinal direction and, when sufficient tissuewas available, the circumferential direction. The original test sectionwidth was measured with digital calipers. The wide flaps of the sampleswere wrapped in 400 grit sandpaper and loaded into the tensile testingapparatus (Instron 5543, Canton, Mass.) via spring-loaded grips.

Application of the testing protocol and acquisition of test data wereachieved using Instron's Merlin software. The testing consisted of aslow ramp at 0.1 mm/sec, 10 precycles from 0.10 to 0.15 N, a 2-minutehold at constant length, then strain to failure at 0.5 mm/sec.Engineering stress (load/initial cross-sectional area), and engineeringstrain ((final—initial length)/initial length) were used to determinethe stress-strain relationship. Ultimate stress and ultimate strain weredefined as the stress or strain at the point when the sample failed.

The average stress-strain relations for freshly harvested, elongated,and control arteries are displayed in FIG. 10. Defining the transitionzone as the nonlinear region separating two approximately linear regionsof different slopes, the transition zone for both the control andelongated arteries ended at about 65% strain, whereas freshly harvestedarteries ended at about 35% strain. The average ultimate stress andstrain in the longitudinal direction was calculated for fresh, controland elongated arteries; only the ultimate stress of control arteries wassignificantly different than elongated and freshly harvested arteries(Table 8). The ultimate stress and strain in the circumferentialdirection were obtained for some fresh and elongated arteries (Table 8).TABLE 8 Mechanical properties for arteries from juvenile donors. FreshlyCultured harvested elongated Cultured control arteries arteries arteriesLongitudinal (n = 9) (n = 4) (n = 5) Ultimate stress (MPa) 1.41 ± 0.131.39 ± 0.21 2.11 ± 0.10*,## Ultimate strain (%) 94.1 ± 7.67  121 ± 12.9 115 ± 9.53 Circumferential (n = 3) (n = 4) Ultimate stress (MPa) 1.98 ±0.46 0.87 ± 0.09 N.A. Ultimate strain (%)  106 ± 3.70 89.9 ± 23.6 N.A.Data were determined from mechanical testing in the axial andcircumferential directions. Data are shown with the standard error ofthe means (SEM).Significance differences were denoted between fresh and control arteries(*), and elongated and control arteries (+).One symbol equals p < 0.05, two equals p < 0.005.N.A. indicates not analyzed.

Tests for Statistical Significance: In cases when the same specimencould be tracked (such as the artery length before and after culture),one-tailed, paired t-tests were used. Otherwise, one-tailed, two-samplet-tests assuming unequal variance were utilized, p<0.05 was consideredsignificant. For figures and tables, one symbol (*) denotes p<0.05,while two symbols (**) denote p<0.005. The ultimate stress and strain inthe circumferential direction were obtained from fresh and elongatedarteries (Table 8). While the ultimate circumferential stress of controlarteries was 2.3 fold. greater than that of elongated arteries, thedifference was not significant (p=0.07). Taken together, these dataindicate that ex vivo cultured vessels retained their viability,structure and function.

Example 5 Ex Vivo Vascular Remodeling

Of the three aspects of arterial remodeling to be investigated (wallthickness, longitudinal length, and internal diameter), the least isknown from in vivo studies about the mechanical factors that regulatethe longitudinal length of a vessel. Therefore, to test whetherlongitudinal stress or strain stimulates vessels to elongate, and tofurther validate the ex vivo perfusion system, four excised porcinecarotid arteries were placed in the ex vivo system and initiallystretched to their in vivo length. Each day the vessels were stretchedan additional ˜10% by sliding the stainless steel tubes shown in FIG.2A. After 9 days in culture, the length of the vessel had increased100%. In contrast, acute stretching of the arteries resulted in ruptureafter about 80% strain (FIG. 98B).

When removed from the system, arteries that had been stretched for 9days shortened (elastic recoil similar to what was observed when vesselswere initially excised from in vivo), but the resulting length was 70±3%greater than the initial length of the freshly isolated arteries priorto stretching, as summarized in FIG. 11. The length of the vessel inline A (the in vivo length) was equal to that in line C (the initiallength in the system). Line B shows the freshly excised length showingelastic recoil from the in vivo length. Line D was the length of thevessel after 9 days in the ex vivo system (showing a 100% increase overthe in vivo length). Line E shows the length of the vessel after removalfrom the ex vivo system, which even after recoil was 70±3% greater thanthe initial length. The length in line D=2.0× that of line C. The lengthin line E=1.7× that of line B. Conditions B and E are unstressed (i.e.,no applied longitudinal loads); while all other conditions involvelongitudinal loading

These data provide evidence that aspects of mechanically inducedvascular remodeling observed in vivo can be reproduced in the ex vivoperfusion system of the invention.

Example 6 Evaluating Performance of Tissue-Engineered Blood Vessels InVivo

To rigorously evaluate the potential utility of ex vivo remodeledarteries for bypass surgery, in vivo studies are being conducted. Exvivo cultured arteries are implanted as autologous interposition leftcarotid artery grafts. The in vivo performance of these grafts withrespect to patency and resistance to intimal hyperplasia are compared toautologous saphenous vein grafts placed interpositionally in the rightcarotid arteries of the same test pigs, and freshly harvested carotidarteries (i.e., no ex vivo culture) placed back in their original donorset. This experimental design allows comparison of the ex vivo remodeledvessels to a positive control (the freshly harvested carotid artery,which is an excellent vascular graft material) and a negative control(the saphenous vein, which is a relatively poor vascular graftmaterial).

Several sets of in vivo studies are conducted, wherein the majordifference between the two sets being the conditioning of the ex vivoremodeled vessels. The first set of studies is designed to test thehypothesis that ex vivo culture of vessels under mechanical conditionssimulating normal physiological loading will result in minimal vascularremodeling, and that the patency of these arteries is approximate thatof freshly harvested vessels. In subsequent sets of experiments, themechanical environment during ex vivo culture is modified to direct theremodeling of the excised vessels.

Example 7 Using the Ex Vivo Perfusion System to Explore the MolecularRegulation of Mechanically Induced Vascular Remodeling

To evaluate the expression of Tenascin-C (TN-C) protein and mRNA inarteries exposed to the different mechanical regimes of the presentinvention, segments of arteries cultured ex vivo are routinely fixed andsectioned to prepare histological sections. Histological sections areimmunostained for TN-C protein following a procedure similar to the oneused to stain for smooth muscle cell α-actin and PCNA (FIGS. 6B, 6E).

The majority of studies show that soluble, extracellular, and matrixfactors regulate TN-C at the transcriptional level, therefore, in situhybridization studies with digoxigenin-labeled TN-C riboprobes are usedto ascertain the regulation of TN-C expression at the mRNA level. Ifmechanically induced changes in TN-C protein levels in the arterial wallare regulated on the mRNA level, the region(s) of the promoterresponsible for mechano-sensitivity are determined using full length anda series of 11 mutated TN-C promoters linked to a CAT reporter gene.These constructs have been previously used to determine to the regionsof the TN-C promoter that regulate TN-C transcription by cultured smoothmuscle cells in response to remodeled type-I collagen (Jones et al., J.Cell Sci. 112(Pt 4):435-445 (1999)).

The TN-C promoter—CAT reporter plasmids are individually incorporatedinto a polylactic acid (PLA) (3 mg PLA/1000 ml chloroform) to give afinal DNA concentration of 14 μg/ml. DNA-polymer emulsions are appliedto the surface of a Dacron mesh, and then desiccated under a laminarflow hood. Plasmid DNA is delivered from an adventitial position bywrapping meshes around isolated arteries prior to their placement in theex vivo organ culture system. Jones and others have used this techniqueto deliver DNA to the arterial wall in vivo.

After a period of ex vivo culture exposed to the desired mechanicalenvironment, the artery is retrieved, and a segment of the vessel isfixed in paraformaldehyde, sectioned and immunostained with antibodiesthat recognize the CAT protein. The remaining segment of the vessel isanalyzed for CAT activity using established techniques. By couplingimmunostaining of histological sections and quantification of CAT enzymeactivity, the spatial distribution and amount of the reporter protein isdetermined. By comparing the CAT expression driven by differentpromoters, the salient region(s) for mechanosensitivity are indicated.Special attention is given to the potential role of a putative shearstress responsive element (GAGACC) 600 base pairs upstream from thetranscriptional start site.

In sum, the controlled, ex vivo vascular remodeling system and method ofthe present invention has been shown to provide a clinically significanttool for the tissue engineering of vascular grafts from small excisedvessels, as demonstrated at the physical and molecular levels, and asoptimized in vivo. Consistent with the principle that tissue-engineeredarteries generated by the present invention more closely resemble thestructure and function of native arteries than arteries constructed fromisolated cells, arteries isolated from juvenile pigs and elongated exvivo were nearly identical to native arteries in terms of structure(both macroscopically and histologically, including endothelial coverageand intricate structural components such as the internal elasticlamina), viability (as measured with the MTT assay and TUNEL analysis),and function (vasoactivity and mechanical properties). Aside fromincreased extensibility at low stress, the biomechanical properties offresh and elongated arteries, notably the ultimate longitudinal andcircumferential stresses and strains, were not significantly differentfrom fresh arteries demonstrating that when the elongated arteries areused as vascular grafts, they are expected to behave in a manner similarto native arteries in terms of mechanical integrity, as well as toprovide clinically relevant patency rates when implanted in vivo.Moreover, ex vivo it is possible to precisely control the mechanicalenvironment while carefully monitoring the resulting growth/remodeling,thereby opening new avenues of research regarding the mechanical stimuliresponsible for specific aspects of remodeling in vivo.

Each and every patent, patent application and publication that is citedin the foregoing specification is herein incorporated by reference inits entirety.

While the foregoing specification has been described with regard tocertain preferred embodiments, and many details have been set forth forthe purpose of illustration, it will be apparent to those skilled in theart that the invention may be subject to various modifications andadditional embodiments, and that certain of the details described hereincan be varied considerably without departing from the spirit and scopeof the invention. Such modifications, equivalent variations andadditional embodiments are also intended to fall within the scope of theappended claims.

1. A method of physically remodeling a small blood vessel, whilemaintaining the viability of the vessel, comprising the steps of:excising the blood vessel from its native site, and subjecting theexcised vessel to a controlled, ex vivo mechanical environment for atime sufficient to remodel the vessel by increasing the diameter,length, or wall thickness of the vessel, or any combination thereof. 2.The method of claim 1, wherein the excised vessel is a small artery or avein.
 3. The method of claim 1, further comprising applying pressure,shear, and strain to the vessel under controlled conditions within themechanical environment, wherein transmural pressure drop regulates wallthickness, longitudinal tension regulates length, and flow-induced shearstress regulates inner diameter of the remodeled vessel.
 4. The methodof claim 3, wherein the mechanical environment is controlled by an exvivo perfusion system.
 5. The method of claim 1, further comprisingusing the remodeled vessel as an arterial graft in vivo.
 6. The methodof claim 1, wherein length of the remodeled vessel is increased at least100% over its native length when excised, and wherein more than 50% ofthe increased length is retained after recoil when the remodeled vesselis removed from the controlled mechanical environment.
 7. A method ofphysically remodeling a small blood vessel to be used in vivo as avessel graft in a patient in need of such a graft, comprising the stepsof: excising the blood vessel from its native site; and subjecting theexcised vessel to a controlled, ex vivo mechanical environment for atime sufficient to increase diameter, length, or wall thickness of thevessel, or any combination thereof; removing the remodeled vessel fromthe ex vivo mechanical environment; and surgically inserting theremodeled vessel in vivo as a vessel graft (artery or vein) into thepatient.
 8. The method of claim 7, wherein the excised vessel is a smallartery or a vein.
 9. The method of claim 7, wherein the excised vesselis autologous to the patient.
 10. The method of claim 7, furthercomprising applying pressure, shear, and strain to the vessel undercontrolled conditions within the mechanical environment, whereintransmural pressure drop regulates wall thickness, longitudinal tensionregulates length, and flow-induced shear stress regulates inner diameterof the remodeled vessel.
 11. The method of claim 7, wherein themechanical environment is controlled by an ex vivo perfusion system. 12.The method of claim 7, wherein length of the remodeled vessel isincreased at least 100% over its native length when excised, and whereinmore than 50% of the increased length is retained after recoil when theremodeled vessel is removed from the controlled mechanical environment.13. An ex vivo perfusion system for exposing a viable, excised bloodvessel to precisely controlled flow and pressure regimes, wherein thesystem comprises: a pump means, which when activated, continuouslypushes fluid through the system; a housing means, comprising amedium-filled chamber, within which chamber the excised vessel ishoused, and the excised vessel is cannulated with two sliding tubes,wherein when activated, the chamber housing the vessel is perfused withcell culture medium supplemented with serum and antibiotics, and whereintemperature, pH, pO₂, pCO₂, and nutrients are maintained at levelssufficient to maintain the viability of the vessel; a reservoir withinwhich the culture medium is pooled, having a gas exchange port, whichpermits gas exchange within the medium; a controller means to controlpressure within the chamber housing the excised blood vessel; an in-lineprobe means to measure and report pressure within the system; a datameasurement means attached to the in-line probe means for digitizing themeasured pressure data; and a computer node attached to the datameasurement means to record, analyze and store the digital data.
 14. Theex vivo perfusion system of claim 13, wherein the system furthercomprises: as the pump means, a pulsatile blood pump, which whenactivated, continuously pushes fluid through the system; as the housingmeans, an enclosed Plexiglas cylinder, which forms the housingcomprising a medium-filled chamber, cannulated on each end, within whichchamber the excised vessel is cannulated with two slidingstainless-steel tubes, wherein the chamber housing the vessel isperfused with cell culture medium supplemented with serum andantibiotics, and wherein temperature, pH, pO₂, pCO₂, and nutrients aremaintained at levels sufficient to maintain the viability of the vessel;a reservoir within which the culture medium is pooled, having a gasexchange port, which permits gas exchange within the medium, before themedium is returned to the pump for circulation within the system; as acontroller, a needle valve controller at either end of the chamber tocontrol pressure within the chamber housing the excised blood vessel; asan in-line probe, at least one in-line probe to measure pressure withinthe system at a rate of approximately 250 times per second, wherein thedata is reported in analog; as a data measurement means, a datameasurement module attached to the in-line probe(s) for digitizing theanalog pressure.
 15. The system of claim 13, wherein the excised vesselis a small artery or a vein.
 16. The system of claim 13, comprising asingle excised blood vessel.
 17. The system of claim 13, comprisingmultiple excised blood vessels run in parallel, each vessel containedwithin its own housing, corresponding chambers and needle valves. 18.The system of claim 13, wherein ports on the Plexiglas cylinder allowthe exchange of medium and nutrients, fluid overflow and air/CO₂discharge.
 19. The system of claim 13, wherein improved control of themechanical environment provides localized intravascular andextravascular pressure measurement and control, which provides real timemonitoring of vessel remodeling.
 20. The system of claim 14, wherein thetwo sliding stainless-steel tubes slide independently of the rest of theunit to control vessel strain.